There are many tragic injuries and illnesses in the world, and one of the most vivid reminders of these is amputees. It is estimated that there are 70,000 major amputations performed annually in the United States alone, and more than 200 million amputations each year in the world. Many of these are major limb amputations of the lower extremities.
In an effort to improve the life of the amputee there has been significant research in developing artificial limbs that look and move more like actual human limbs. Advances in robotics, biomechanics, composites and computers have made great strides in realistic artificial limbs. Prosthesis is defined as an artificial extension that replaces a missing body part. An artificial limb is a type of prosthesis that replaces the missing body part, such as an arm and leg.
There are four main types of artificial limbs, namely transtibial, transfemoral, transradial, and transhumeral, wherein the prosthesis depends on what part of the limb is missing. The new plastics and composites which include carbon fiber provide for greater strength as well as greater design options using computer aided design.
Artificial limbs are typically manufactured by a process that that involves; Measurement of the stump area Body measurements to determine the artificial limb size; Creation of a model of the stump; Formation of a thermoplastic sheet around the model for fitting; Formation of a permanent socket; Formation of the plastic or composite parts of the artificial limb; Creation of metal parts of the artificial limb; and, Assembly of entire limb.
For illustrative purposes, a prosthetic leg includes a number of different elements that allow for the proper motion and to distribute the various forces. The first element is typically a liner that resides on the residual limb and attaches through a prosthetic lock to the other parts of the prosthesis. The liner is a soft stretchy material that protects the limb and acts as an interface between the socket and the limb.
The residual limb and liner are then coupled to a hard socket. The socket is specially made to fit the specific user and comes in a variety of materials. The socket is custom made to create a better fit between the residual leg and the artificial limb. The socket can be created by taking a plaster cast of the stump and then making a mold from the plaster cast, however other methods such as laser guided measuring can be utilized.
The additional elements for the leg prosthetic depend upon the prosthetic. For example, a transfemoral prosthesis has a prosthetic knee joint connected to the socket to allow for the knee flexure. A prosthetic foot is the last element and is designed with properties that allow for stability and movement to enable walking. In between the main elements of the prosthetic legs are smaller adapters that connect the main parts together and help in the proper alignment of the prosthesis. Many artificial limbs are attached to the stump of the amputee by belts and cuffs or by suction.
Current prosthetic alignment techniques are primarily based on visual estimation, and while there are some devices for static alignment, their tuning is modified after the dynamic alignment is done since leg loads are changed when moving the prosthetic. The muscle forces exerted by the human body on the prosthetic cannot be accurately taken in consideration when doing static alignment only, and dynamic alignment is typically the most important part of the process of prosthetics fitting.
The alignment practices for prosthetic devices vary, depending on the design characteristics of the prosthetic, the location of the artificial limb and the individual. However, the goal of alignment is to establish a stable and safe prosthetic limb. By way of illustration, the prosthetic knee components vary in design, shape, size, resistance to buckling, and swing characteristics.
Typically, the knee is aligned to center posterior to what is considered the reference or “weight bearing” line. The posterior placement creates an extensor moment at the knee, generally resulting in more stability. The alignment stability can be evaluated in the sagittal plane with respect to three commonly used reference lines, namely the European line, the trochanter, knee, and ankle (TKA) line, and the medial, knee, and ankle (MKA) line. The European line projects vertically along the lateral aspect of the limb, passing through the center of the socket brim and bisecting the horizontal length of the foot. The TKA line extends through the head of the greater trochanter, center of the knee, and the center of the ankle. The MKA line is the line from the medial center of the socket to the center of the knee to the center of the ankle.
The older static alignment processing used a vertical reference line formed by dropping a plum bob from the trochanter and measuring the position of the knee center and the ankle from that line. Newer alignment tools form a similar reference line by projecting a vertical laser line up from a force plate onto the patient. This projected laser line may indicate either the center of pressure or be translated to a landmark on the patient, such as the trochanter, thus applying a visual demarcation similar to a plumb line. Regardless of the initial static alignment, there is an ongoing dynamic alignment that addresses the actual load related changes required for an optimal fit.
As noted herein, most rehabilitation clinics and medical research institutes use specialized alignment protocols, based on ‘cause’ related classifications of movement disorders. There is currently no known objective way to get optimal alignment, as it is substantially based on subjective experience of the prosthetist. The results vary greatly between different clinics and rehabilitation hospitals, and often a patient that had his prosthetic limb aligned by one prosthetist will typically have to undergo a second and third repetition of the alignment process for optimal performance. At the present time, the state of the art does not seem to have a known system or method available for dynamically aligning accurately and objectively prosthetics such as prosthetic legs.
Modern prosthetic technologies are more complex but offer a greater span of dynamic function to the user when aligned properly. The C-Leg microprocessor prosthetics try to optimize function and take advantage of the unique characteristics afforded by microprocessor and software control.
For illustrative purposes, examples herein are provided with respect to walking. The term gait refers to the manner of walking, wherein a full gait cycle is defined as the time interval between two successive occurrences of one of the repetitive events of walking. There are typically seven identified major components to a full gait cycle, namely initial contact, opposite toe-off, midstance, heel rise, opposite initial contact, toe-off, feet adjacent, and tibia vertical.
The starting point in the gait cycle is typically at initial contact. For example, if the left foot is the starting point or reference, the cycle continues until the left foot makes initial contact again. The distance covered during this cycle is called a stride length, wherein a step length is one half the stride length.
The gait cycle components can be sub-divided into two phases, namely the stance phase and the swing phase. The stance phase occurs when the foot is making contact with the ground and involves the first four components of the gait cycle, namely initial contact, opposite toe-off, midstance and heel rise. The swing phase is when the foot is moving forward through the air and consists of the remaining components of the gait cycle, namely opposite initial contact, toe-off, feet adjacent, and tibia vertical.
Initial contact refers to heel contact, wherein the major role of the lower extremity is to absorb the impact forces created when the foot strikes the ground. The heel pad along with control of the ankle as the foot moves from the heel to the forefoot allows the absorption of the forces. On contact, the hip extensors, gluteus maximus, and hamstrings create internal extensor movement at the hip, and these muscles concentrically contract and propagate an extension force at the hip joint. Simultaneously, the knee undergoes an internal flexor moment secondary to the contraction of the hamstrings to block hyperextension at the end of the swing phase. The ankle is usually kept in the neutral position on initial contact in preparation for the next phase.
The loading response refers to the double support period between the initial contact and opposite toe-off components. The foot is lowered to the ground by means of plantarflexion of the ankle, which is simultaneously resisted by dorsiflexion produced by tibialis anterior. These actions maintain control of the foot and allow for a gentle lowering to the ground. At this point, the center of gravity is at its lowest point in the gait cycle.
After the loading response, opposite toe-off is the next component of the gait cycle and is the beginning of midstance and the first period of single support. The forefoot impacts the ground at about the same time opposite toe-off occurs. The hips move steadily through extension with power, while the knee generates an external flexor movement. The quadriceps muscles eccentrically contract, absorbing energy and allowing the knee to act like a spring. The direction of the ankles shifts from plantarflexion to dorsiflexion when the tibia passes over the stationary foot.
The term Midstance is defined as the period between opposite toe-off and heel rise. It signifies the moment when the swing-phase leg passes the stance-phase leg. During the period, the hip begins to lose its extensor movement with a decline in contraction of the gluteus maximus and hamstrings. The knee shifts its motion from flexion to extension and at the same time generates power. As the tibia moves forward over the ankle due to the inertia created by the trunk, it undergoes external rotation concomitant with forefoot supination. The ankle continues to shift from plantarflexion to dorsiflexion with the triceps surae muscle contracting eccentrically. The speed at which the center of mass of the body moves over the supporting stance-phase limb is regulated by the power created during plantarflexion of the ankle.
Heel rise is the next component of the gait cycle and represents the period when the heel begins to lift from the walking surface. A progressive internal flexor moment is created at the hip, while an internal flexor knee moment is initiated when the quadriceps muscles stops contracting before heel rise. The knee action occurs because the upper body moves faster that the tibia and because the triceps surae retards the forward motion of the tibia while the femur steadily moves forward. These motions create an external extensor moment opposed by an internal flexor moment at the knee. The ankle has an internal dorsiflexor moment as the soleus and the gastrocnemius begin to progressively contract.
The Opposite initial contact starts pre-swing with the start of opposite initial contact, the hip and knee begin to flex while the ankle is plantarflexing. The body now pivots on the forefoot instead of the ankle, which creates more power, and the triceps surae and other secondary ankle plantarflexors create a corresponding internal plantarflexor moment in response to the external dorsiflexor moment. These muscles use an eccentric contraction. The triceps surae is used to impede the body's momentum instead of launching it forward and allows favorable ankle stabilization and a decline in the amount of fall by the body's center of gravity. In addition, the adductor longus muscle acts as the primary hip flexor in this phase, and the rectus femoris muscle contracts eccentrically to stabilize knee flexion. All of these actions assist with forward acceleration of the leg into the swing phase.
The next component, toe-off, represents the end of the stance phase and the beginning of the swing phase. Muscle contraction changes from eccentric in stance phase to concentric in swing phase. Toe-off occurs at around 60% point of the gait cycle. An internal flexor moment occurs at the hip secondary to inertial forces and contraction of the adductor longus and iliopsoas muscles. The rectus femoris muscle contracts to prevent excessive knee flexion and the internal plantarflexion moment loses power at the ankle as the toe leaves the ground.
Following the toe-off component, the feet adjacent component is the next stage of the swing phase. Considerable power is generated at the hip by the rectus femoris, adductor, and iliopsoas muscles to move the leg forward through the swing phase. Eccentric contraction of the quadriceps continues throughout the first half of the swing phase to regulate the rate and extent of knee flexion. Some of the kinetic energy created through contraction and inertia is transferred to the trunk as the swing leg is decelerated at the end of this phase.
The final component of the gait cycle is tibia vertical, which is represented by the tibia of the swinging leg becoming vertical. This is the period between midswing and terminal swing. In this phase, the knee extends in preparation for the beginning of the stance phase. This extension is accomplished through two mechanisms, namely concentric contraction of the hip extensors which causes a posterior rotation at the thigh, and inertia created at the foot and shank which allows it to continue forward. Eccentric contractions of the hamstrings gradually decelerate the foot and shank until the knee arrives at an extended position. At this point, the swing phase leg not only is prepared for the next stance phase but also helped with trunk acceleration. As previously noted, some of the kinetic energy created during the swing phase is transferred to the trunk upon deceleration.
People who have undergone amputations generally incorporate different muscles and adaptive strategies to ensure a smooth and well-coordinated gait pattern. One of the underlying attributes in gait analyzing an amputee's gait is to try to use the least amount of energy to cover the greatest distance. Several factors are typically considered when thinking about the energy costs of prosthetic ambulation. One is the actual metabolic costs which is the peak exercise oxygen consumption [VO2] in mL/kg/m) of the person who has undergone amputation compared to that of intact people. Increased metabolic cost for persons who have undergone amputations means the gait is inefficient compared to that of healthy intact persons, who require less endurance for any given distance.
A pathologic gait is inefficient and usually requires considerably more energy than a normal gait. Patients may adopt many kinds of abnormal movements to minimize their energy usage, which can be categorized into those involving energy transfers and those involving movements that minimize the displacement of the center of gravity. The optimizations or determinants of gait can be generally considered to be Pelvic rotation, Pelvic obliquity, Knee flexion in the stance phase, Ankle mechanisms, Foot mechanisms, and Lateral displacement of the body. In a general sense, these elements can be combined to create a smoother gait and reduced energy expenditure by minimizing the downward and lateral motion of the center of gravity.
Patients with various amputations have adopted strategies for minimizing their energy consumption in ambulation. In persons who have undergone TT (transtibial) amputations, the timing and magnitude of the muscular work patterns in the intact limb are correlated with a normal gait. However, the prosthetic limb must make up for the energy absorption of the quadriceps and triceps surae muscles and for the eccentric power generation of the triceps surae.
During the stance phase, the prosthetic limb performs about half of the work of normal muscle. Energy absorption by the knee extensors and energy generation by the prosthetic foot are substantially reduced. To offset the loss of power from triceps surae with roll-off, the person who has undergone amputation changes the biomechanics of both the prosthetic limb in the stance phase and the intact limb during the swing phase. In the prosthetic limb, the primary energy absorbers and generators shift to the hip extensors during the stance phase. Also, during swing phase, the muscular work components substantially increase in the intact limb. The excess mechanical work is ultimately transmitted to the trunk during terminal swing-phase deceleration. The increase in the forward momentum of the trunk compensates for the loss of power generation with the prosthetic foot.
The person who has undergone TF (transfemoral) amputation must deal with the loss of the foot, ankle, and the knee. The biggest apprehension of the person who has undergone TF amputation is the prevention of knee buckling. Besides the actual prosthetic hardware and knee alignment to add stability, the biomechanics of the gait are changed to provide additional stability. The person who has undergone TF amputation does not allow knee flexion in the first 30-40% of the stance phase. This limitation minimizes the likelihood of knee buckling. In addition, the hip extensors help maintain hip extension through closed kinetic chain mechanisms.
In the opposite initial contact part of the gait cycle, the ankle plantarflexors, particularly the triceps surae and the hip flexors, contract to generate power for the acceleration of the leg forward into the swing phase. Although the prosthetic limb is only a fraction of its normal mass with a TF amputation, the hip flexors must generate the same power as a normal limb. The intact limb adjusts to compensate for the prosthetic limb in these cases.
During the stance phase in the intact limb, generated energy is augmented by the hip extensors and the ankle plantarflexors. In this way, person who has undergone amputation tries to offset the loss of power from the triceps surae in the prosthetic limb. During the swing phase, the biomechanics of the person who has undergone TF amputation mimics those of normal gait, including the energy-absorbing function of the quadriceps performed through the prosthetic hydraulic knee unit.
The common gait deviations of TT and TF prosthetic gait are briefly addressed in terms of the types of amputation and the times at which the deviations occur in the gait cycle.
Stance-phase problems can occur in the gait of individuals with a TT prosthetic. Inappropriate knee flexion can occur in the early stance phase, causing knee instability. Several problems could arise from this flexion, including excessive ankle dorsiflexion, socket flexion, and posterior foot placement. Knee hyperextension could also occur in the early stance phase, emanating from ankle plantarflexion or socket extension, weak knee extensors, anterior foot placement, or inadequate prosthetic foot selection.
Mediolateral knee thrust can also be observed in the stance phase. This is usually derived from inadequate side-to-side placement of the foot, excessive angulation of the socket, or wide mediolateral proximal socket dimensions that cause decreased knee control. If an individual who has undergone a TT amputation is noted to have a foot slap in his or her gait, it may be a result of excessive socket flexion or foot dorsiflexion, the uneven placement of the foot, or a deficient heel height for proper prosthetic alignment. Excessive forward progression of the tibia, or a drop-off gait can be caused by impaired rollover, shortening of the contralateral step length and swing time, and delayed heel-off.
External rotation can occur at two different phases of the gait cycle: heel strike or late stance. If external rotation occurs during heel strike, the etiology could be a solid ankle cushion heel (SACH) durometer that is too dense, an articulated foot plantarflexion bumper that is too hard, or misplacement of the suspension cuff-retention points. If external rotation occurs in late stance, it can be caused by inadequate excessive anterior placement of the foot, excessive foot plantarflexion, or excessive hardness of the forefoot.
Early heel rise could result from inadequate placement of the foot (posterior), flexion contracture in the hip or the knee that was not accounted for in fitting the prosthetic, or excessive softness of the forefoot. Contralateral early heel rise or vaulting is a pathologic gait that allows clearance of the prosthetic limb with decreased hip and knee flexion. Vaulting compensates for a prosthesis that is too long, inadequate suspension of the prosthesis, or a learned gait pattern.
Fewer gait problems are involved with the swing phase than with the stance phase. The objective of the swing phase is forward advancement of the non-weightbearing limb. When prosthetic limb clearance is poor, the gait becomes pathologic. Most predicaments occur because of poor suspension, a prosthesis that is too long, insufficient prosthetic knee flexion, or inadequate transfer of power from the residual limb to the prosthesis that decreases or delays knee flexion. A coordinated, smooth, swing phase is facilitated by energy-efficient limb clearance, which is enabled by synchronized motion at the hip and knee joints and by total joint displacement.
Foot drag is one of the most common problems of swing phase. It is usually caused by inadequate suspension of the prosthesis, a prosthesis that is too long, or lower-limb weakness in the hip abductors or ankle plantarflexors on the contralateral side. Any abnormal limb rotation is observed during the gait trial is usually caused by insufficient suspension of the prosthesis, misplacement of the suspension cuff-retention points, or overshooting of the hip or knee flexors to evade foot drop.
Limited knee extension or flexion problems can always be traced back to mechanical contractures, problems with the suspension, or problems with the knee joint in relation to a thigh corset.
The gait deviations in persons who have undergone TF amputations differ from those of people who have undergone TT amputations in a couple ways. As alluded to previously, knee flexion in the stance phase is one of the most common problems related to gait instability with TF amputations. If a patient with a TF amputation is concerned about putting weight on the prosthetic leg because the knee flexion moment creates instability, an inefficient gait pattern results. For these patients, one of several unique models of prosthetic knees can be prescribed based on his or her individual needs.
Additional etiologies of knee flexion in stance are a hard SACH durometer, excessive foot dorsiflexion, excessive socket flexion, weak hip extensors, or decreased weightbearing capability. Prolonged knee extension in the stance phase is another problem that can occur with TF amputations. This extension can result in shortening of the contralateral step and an increase in the vertical displacement of the center of gravity.
If lateral hip thrust is the problem, immediate attention should be given to the wide dimensions of the mediolateral proximal socket that affect the stability of the hip. If the socket fits well, the patient could have weak hip abductors, or the hip adductors might not have been reattached at the time of surgery.
The most prevalent gait abnormality with TF amputations is ipsilateral trunk bending in the stance phase. Similar to the compensation for Trendelenburg gait, this abnormality could indicate weak hip abductors on the ipsilateral side or an inappropriately short prosthesis. Occasionally, a person who has undergone a TF amputation can have an awkward downward movement of the upper body over the prosthesis, especially during fast walking. This is referred to as a drop-off gait in the late stance. The foot of the prosthesis should be checked for excessive dorsiflexion whenever external rotation of the leg occurs, either during heel strike or in the late stance. The foot should be examined for excessively hard materials. As with TT amputations, TF amputations can be associated with some of the same problems in the stance phase; the etiologies of the abnormal biomechanics are similar.
In people who have undergone TF amputations, gait abnormalities in swing phase are limited in number. Stiff-knee gait patterns can be the consequence of excessive knee stability in the joint that makes the creation of a flexion moment at the knee difficult. Circumduction, or the swing of the limb in a wide arc, usually indicates inadequate suspension or excessive length of the prosthesis. Abnormal axis rotation at the knee that results in a whipping motion is usually due to incorrect alignment of the prosthesis at the knee.
Thus, the dynamic alignment process requires not only a keen perception by the alignment personnel, but considerable experience in evaluation the numerous factors. The alignment of every person is different and there are no quantitative guidelines for the subjective optimization.
What is needed, therefore, are techniques for real time dynamic alignment of prosthetics utilizing a smart motion blending algorithms and a blending engine to create a single objectified protocol for the alignment process.